Method and apparatus for enhanced X-ray computing arrays

ABSTRACT

An x-ray imaging system utilizes enhanced computing arrays. A plurality of x-ray illumination source positions are utilized to produce x-ray radiation at each of the x-ray illumination source positions and to project x-ray radiation towards an object. A detector detects x-ray radiation from the object and transmits detector images for each of the illumination source positions. A memory buffer stores the detector images from the detector. A graphics processing unit formats the detector images and constructs a complete frame data set with the detector images for each of the illumination source positions. Another graphics processing unit receives the complete frame data set and performs image reconstruction on the complete frame data set.

RELATED U.S. APPLICATION

This application claims priority to the U.S. provisional patentapplication Ser. No. 62/065,585, entitled “Method and Apparatus forEnhanced GPU Computing Device Arrays,” with filing date Oct. 17, 2014which is incorporated by reference herein in its entirety.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH AND DEVELOPMENT

The U.S. Government may have a paid-up license in this invention and theright in limited circumstances to require the patent owner to licenseothers on reasonable terms as provided for by the terms of Grant Nos.5R44EB015910-03 awarded by the National Institute of Health (NIH).

FIELD OF THE INVENTION

The present invention relates to imaging and display systems. Moreparticularly, the present invention pertains to a method and apparatusfor enhanced computing device arrays in medical x-ray imaging anddisplay.

BACKGROUND

X-ray imaging involving tomosynthesis or computed tomography can involveprocessing of large quantities of image data. The object is viewed frommultiple angles and image data from each angle is processed and combinedwith tomosynthesis or computed tomography techniques. Imaging a movingobject particularly a fast moving object such as a beating heart imposesfurther challenges on the imaging system. In order to present an imageto the user with smooth or fluid motion, a frame rate of 7.5 or 15 oreven 30 frames per second may be required, further increasing the amountof image data processed. When the x-ray imaging system is deployed in asurgical environment, e.g. cardiac catheterization, angioplasty orablation, the image processing must be accomplished in real-time,introducing further performance and speed requirements to achieveprocessing in allowable time budgets.

A scanning electron beam x-ray imaging system can utilize an x-raysource array with multiple illumination focal spot positions e.g.hundreds or thousands of illumination focal spot positions. Eachillumination focal spot position can generate a separate detector imageand the performance and speed requirements imposed on the imaging systemcan be enormous particularly at high frame rates for real-timeapplications.

Graphics processing units (GPUs) can be utilized in high performancex-ray imaging systems. Use of GPUs can be beneficial over imagingsystems based on central processing units (CPUs) because of the GPUsability to process imaging data in parallel rather than the serialnature of CPUs. Performance of GPU-based x-ray imaging systems can beenhanced by using multiple GPUs. But when incoming imaging data isstreamed to these GPU devices at a continuous rate, severe performancelosses by the individual GPUs can occur. The performance losses canresult in the use of additional GPUs in the imaging system increasingthe cost and complexity of the system or can prevent the imaging systemfrom achieving desired image quality and high frame rates.

What is needed is a cost-effective imaging and display system capable ofprocessing and producing high quality images at high frame rates in areal-time environment. Furthermore, the imaging and display systemshould have speed and performance characteristics allowing use withmultiple illumination focal spot positions and x-ray source arrays.

SUMMARY

The present invention pertains to a method and apparatus for x-rayimaging system utilizing enhanced computing arrays. A plurality of x-rayillumination source positions are utilized to produce x-ray radiation ateach of the x-ray illumination source positions and to project x-rayradiation towards an object. A detector detects x-ray radiation from theobject and transmits detector images for each of the illumination sourcepositions. A memory buffer stores the detector images from the detector.A graphics processing unit formats the detector images and constructs acomplete frame data set with the detector images for each of theillumination source positions. Another graphics processing unit receivesthe complete frame data set and performs image reconstruction on thecomplete frame data set. A scanning beam x-ray source can be utilized. Adirect memory access unit can be utilized with one or more of thegraphics processing units. A bus can be utilized to transfer thecomplete frame data set between graphics processing units in a singleburst operation, which can be a PCI-E bus. Yet another graphicsprocessing unit that receives a complete frame data set can be utilizedto perform image reconstruction.

Detector images can be transferred from the detector at a rate of atleast 1 MHz. The graphics processing units can utilize blocks of workerthreads capable of performing independent calculations and memoryoperations. A complete frame data set can be produced at least onceevery 130 milliseconds. The complete frame data set can be transferredat a rate of 30 Hz. The complete frame data set can be transferred toone of the graphics processing units in a single burst operation.Detector images that are dropped from said complete frame data set canbe identified. Image data for each of the illumination source positionscan be aggregated to form the detector images. The first graphicsprocessing unit can transpose the complete frame data set.

These and other objects and advantages of the various embodiments of thepresent invention will be recognized by those of ordinary skill in theart after reading the following detailed description of the embodimentsthat are illustrated in the various drawing figures.

BRIEF DESCRIPTION OF THE DRAWINGS

The present invention is illustrated by way of example, and not by wayof limitation, in the figures of the accompanying drawings and in whichlike reference numerals refer to similar elements.

FIG. 1 is a diagram showing a portion of an exemplary scanning-beamx-ray imaging system of one embodiment of the present invention.

FIG. 2 is a diagram showing an exemplary scanning-beam x-ray imagingsystem of one embodiment of the present invention with a collimationgrid having a plurality of focal points to obtain stereoscopic x-rayimages.

FIG. 3 is a diagram showing a single x-ray beam and generation ofinformation for 5 image pixels.

FIG. 4 is a diagram showing the sequential positions of the axes ofx-ray micro-beams from x-ray pencil beams emanating from fiveconsecutive apertures illuminating a single image pixel.

FIG. 5 is a diagram showing X-ray pencil beams from a first aperture anda second aperture passing through an object plane a distance S₀ from theapertures and passing through a plane a distance S₁ from the apertures.

FIG. 6 is a diagram showing the divergence of a single x-ray pencil beamfrom an aperture to the multi-detector array and the intersection withan object at an object plane.

FIG. 7 is a diagram showing the axes of all of the x-ray micro-beamsfrom all of the apertures that intersect a single image pixel in anobject plane as they travel to the multi-detector array.

FIG. 8 is a diagram illustrating one embodiment of the present inventionin which a region of interest has been defined within the patientvolume.

FIG. 9 is a table displaying contrast-to-noise improvement for differentiodine concentrations for procedures in which iodine has been used as acontrast agent.

FIG. 10 is a plot illustrating a manner of equalization filtration ofone embodiment of the present invention.

FIG. 11 is a diagram showing an overall architecture of one embodimentof the present invention.

DETAILED DESCRIPTION

Reference will now be made in detail to embodiments of the presentinvention, examples of which are illustrated in the accompanyingdrawings. While the invention will be described in conjunction withthese embodiments, it will be understood that they are not intended tolimit the invention to these embodiments. On the contrary, the inventionis intended to cover alternatives, modifications and equivalents, whichmay be included within the spirit and scope of the invention as definedby the appended claims. Furthermore, in the following detaileddescription of embodiments of the present invention, numerous specificdetails are set forth in order to provide a thorough understanding ofthe present invention. However, it will be recognized by one of ordinaryskill in the art that the present invention may be practiced withoutthese specific details. In other instances, well-known methods,procedures, components, and circuits have not been described in detailas not to unnecessarily obscure aspects of the embodiments of thepresent invention.

FIG. 1 is a diagram showing an exemplary imaging system of oneembodiment of the present invention. The imaging system can comprisescanning x-ray source 10. X-ray source 10 can be the x-ray sourcedescribed more fully in U.S. Pat. Nos. 5,682,412 and 6,198,802, entitled“X-Ray Source” and Scanning Beam X-Ray Source and Assembly”respectively, all of which are hereby incorporated herein by referencein their entirety. An imaging system is also further disclosed in U.S.Pat. Nos. 5,651,047, 6,183,139, 6,198,802 and 6,234,671, entitled“Maneuverable and Locateable Catheters,” X-Ray Scanning Method andApparatus,” “Scanning Beam X-Ray Source and Assembly,” “X-Ray Systemwith Scanning Beam X-Ray Source Below Object Table,” respectively, allof which are incorporated herein by reference in their entirety.

Use of scanning x-ray source 10 allows for utilization of a reversegeometry configuration for the imaging system. In a reverse geometryconfiguration, a smaller detector can be used whereas a point sourcerequires a much larger detector. The area of the detector can be 2, 3,4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19 or 20 percentof the area of the maximum field of view for given source configurationand detector distance from patient or any percentage in between suchpercentages or any range of percentages in between such percentages. Asmaller detector allows greater flexibility in positioning the detectorwith respect to the patient.

In a reverse geometry configuration, the detector can also be locatedfurther away from the patient than a detector with a point x-ray source.With a point x-ray source, the size of the detector required for a givenmaximum field of view size increases with the distance of the detectorfrom the patient. The already large detector required with a point x-raysource becomes even larger with increasing distance. With a scanningx-ray source in a reverse geometry configuration, the size of thedetector required for a given maximum field of view size decreases withthe distance of the detector from the patient. Thus, the detector forscanning x-ray source 10 can be located with a distance from the patientof 1.3 m, 1.4 m, 1.5 m, 1.6 m 1.7 m, 1.8 m, 1.9 m, 2 m, 2.1 m 2.2 m, 2.3m, 2.4 m, 2.5 m or any distance in between such distances or any rangeof distances in between such distances.

X-ray source 10 can comprise deflection yoke 20 under the control ofscan generator 30. Deflection yoke 20 can comprise one or more magneticfocus or deflection coils. The magnetic focus or deflection coils can bemade with insulated electrical wire wound around a core. The core can beferrite, steel, iron or other magnetic alloy. An electron beam 40generated within x-ray source 10 can be scanned across target 50 withinx-ray source 10 in a predetermined pattern. Target 50 can be a groundedanode target. The predetermined pattern can be a raster scan pattern, aserpentine (or “S” shaped) pattern, a spiral pattern, a random pattern,a gaussian distribution pattern centered on a predetermined point of thetarget, or such other pattern as may be useful to the task at hand. Theserpentine (or “S” shaped) pattern can eliminate the need in a rasterscan pattern for horizontal “fly back.”

As electron beam 40 strikes target 50 at focal spot 60, a cascade ofx-rays 70 is emitted and travel outside of x-ray source 10 toward theobject 80 to be imaged. To optimize system performance of the presentembodiment, a cone of x-ray photons can be generated that will divergein a manner that will just cover the multi-detector array 110. Thedetector including multi-detector array 110 is further described in U.S.Pat. No. 5,808,306, entitled “X-ray Detector,” which is herebyincorporated herein by reference in its entirety.

This divergence can be accomplished by placing a collimating assemblybetween the target 50 of the scanning x-ray source 10 and themulti-detector array 110, and can be between the target 50 and theobject to be imaged. The collimating assembly can be a collimation grid90, comprising a grid of x-ray transmissive apertures 140. Collimationgrid 90 can be designed to permit passage of only those x-ray pencilbeams 100 whose axes lie in a path that directly interceptsmulti-detector array 110. Collimation grid 90 can be stationary withrespect to multi-detector array 110 while the system is in operation.Thus, as electron beam 40 is scanned across target 50, at any givenmoment there is only a single x-ray pencil beam 100 which passes throughobject 80 to multi-detector array 110.

The output of multi-detector array 110 can be processed and displayed onmonitor 120 as luminance values. Image processing techniques can be usedto produce a computer driven image on an appropriate display orphotographic or other medium.

The imaging system disclosed herein is a low exposure system in that itcan expose the patient at a rate of about 0.09 to 0.33 R/min with a 30frame/sec refresh rate measured at the entrance to the patient, which inother systems under the same conditions can typically be between 2.0 to2.8 R/min. Whole body exposure with a 30 frame/sec refresh rate can belower as well.

Collimation grid 90 can comprise an array of apertures 140, the axes ofeach, are oriented or pointed toward multi-detector array 110. That isto say that the axes of apertures within the collimation grid 90 are notparallel to each other and form an acute to the line perpendicular tothe output face 260 of the collimation grid 90. For example, acollimation grid for chest imaging can comprise apertures forming anangle with a line perpendicular to the output face 260 of thecollimation grid 90 of between 0 degree at the center of the collimationgrid 90 to as much as 20 degrees at the edge of the grid 90. A breastimaging application on the other hand can have a collimation grid 90comprising apertures forming an angle with a line perpendicular to theoutput face 260 ranging to 45 degrees at the edge of the grid. Thus, adifferent collimation grid 90 can be selected and inserted for use indifferent imaging applications.

The number of apertures 140 in collimation grid 90 can correspond to thenumber of image pixels to be generated. For example, 500 by 500 to 1024by 1024. Alternatively, the image pixel to aperture ratio can beincreased, i.e., fewer apertures than image pixels may be used, inconjunction with the technique of “sub-sampling.” The system spatialresolution can be determined, in part, by the pitch of the apertures incollimation grid 90. The precise number of apertures suggested above isillustrative only, and is not intended in any way to be limiting.

The x-ray absorbent portion of preferred collimation grid 90 can bedesigned to absorb errant x-rays so that they do not illuminate object80. This can be accomplished by fabricating collimation grid 90 withsufficient thickness so that the x-ray radiation passing through anaperture 140 towards the multi-detector array 110 is substantiallygreater than the cumulative x-ray radiation passing through the x-rayabsorbent portion in all directions other than toward multi-detectorarray 110. Such errant x-rays would provide the object 80 and attendingstaff with x-ray dosage but contribute no meaningful information to theimage.

Square apertures 140 can be used and can be 0.0381 cm (0.015 in) by0.0381 cm in dimension while round apertures can be 0.015 in (0.038 cm)in diameter. Both square and round apertures can yield a cross sectionalarea at multi-detector 110 that can be about 1/100 the cross sectionalarea of other detectors. The cross sectional area of the face of themulti-detector array 110 can be much smaller than in other conventionalsystems. As a result, x-rays scattered at the object miss themulti-detector array and do not tend to fog the image as they do inother conventional systems which typically utilize relatively largesurface area detectors.

FIG. 2 is a diagram showing an exemplary scanning-beam x-ray imagingsystem of one embodiment of the present invention with a collimationgrid having a plurality of focal points to obtain stereoscopic x-rayimages. The axes 103 of the x-ray pencil beams 100, corresponding to theaperture axes of every other row of apertures 140 a in grid 90 can bepointed at focal point F1 at the center of multi-detector array 92 andthe aperture axes of the remaining apertures are pointed at focal pointF2 at the center of multi-detector array 93. One can scan the aperturesin a raster or serpentine pattern and create a “line” of data from thefirst multi-detector array, and a line of data from the secondmulti-detector array. Repeating this, it is possible to build up twocomplete images, as seen from two distinct angles and thereby displaythem with conventional stereoscopic imaging display systems to provide astereoscopic x-ray image.

Apertures 140 a, 140 b can diverge from a common first aperture 140 toform a “V” as shown providing separate paths along the “legs” of the “V”for x-ray pencil beams 100. There is no requirement, however, thatapertures 140 a, 140 b diverge from a common aperture as shown, but anadvantage of the “V”-shaped aperture where the x-rays enter at thecommon aperture or apex of the “V’ is that both multi-detector arrays 92and 93 can be illuminated simultaneously, the “V” acting as an x-raysplitter with some of the x-rays going to multi-detector array 92 andsome to multi-detector array 93. This can decrease by 50% the powerrequired for the beam current.

To achieve resolutions of several line pairs per millimeter or more atthe object plane, the spatial resolution limit in some reverse-geometrysystems is in large part determined by the size of the singlenon-segmented detector. Generally speaking, a small non-segmenteddetector can provide high spatial resolution while a large non-segmenteddetector provides high collection efficiency. This trade-off can be aproblem in developing low dosage x-ray imaging systems.

When such a detector is small to increase resolution, a large proportionof the x-rays emitted by target 50 are unused by the single detectoreven when a collimator grid 90 is used. This is, in fact, how industrialreverse-geometry scanning-beam x-ray inspection systems are designed,where dose is usually not a consideration. Accordingly, while one candecrease the size of a detector by placing, for example, a lead washerin front of the single detector and thereby increase spatial resolution,the x-ray intensity and/or exposure time would have to be increased tomaintain contrast resolution.

By fabricating a multi-detector array having a large area subdividedinto multiple smaller detector array elements, a large capture area isachieved, while simultaneously through image reconstruction techniquesretaining an image resolution that is comparable to the size of a singlesmall detector element without increasing x-ray intensity and/orexposure time.

The resolution defined by the individual detector elements is maintainedby distributing and summing the outputs from the individual detectorelements into a memory buffer in which each address, i.e., image pixel,corresponds to a specific location in the object plane 280. As anelectron beam 40 is moved discretely across the target 50, illuminatingthe area behind selected apertures 140 of the collimation grid 90, theaddress, to which the output of a given individual detector element isadded, changes. The imaging geometry is shown in FIGS. 3 and 4. In FIG.3, a single x-ray beam 100 is shown along with how it generatesinformation for 5 image pixels. Effectively, the single x-ray pencilbeam 100 emanating from individual aperture 141 is divided into x-raymicro-beams, the number of x-ray micro-beams created corresponding tothe number of individual detector elements 160 which comprise themulti-detector array 110. In the case shown in FIG. 3, the axes of fivex-ray micro-beams 141 a, 141 b, 141 c, 141 d and 141 e are shown. InFIG. 4, the sequential positions of the axes of the x-ray micro-beamsfrom x-ray pencil beams 100 emanating from five consecutive apertures141 through 145 illuminating a single image pixel (“IP”) are shown. Theoutputs from the five individual detector elements 161, 162, 163, 164and 165 receiving the x-ray flux from the five x-ray micro-beams, 145 a,144 b, 143 c, 142 d and 141 e respectfully, are added together toprovide the luminance for the single pixel IP.

Stated differently, the output for each of the individual detectorelements 160 is stored for later summation in an image buffer, at amemory address that corresponds to a very small specific region in theobject plane 280, e.g., a single image pixel.

Accordingly, in one embodiment the memory storage address for the outputof each individual detector element 160 changes with the position of thescanning x-ray beam 40 in an ordered fashion such that each memoryaddress contains the sum of the radiation passing through a specificimage pixel or spot in the object plane 280. In this way the spatialresolution of the system is determined by the size of a singleindividual detector element 160, while the contrast resolution of thesystem is determined by the area of all of the individual detectorelements comprising the multi-detector array 110.

An additional benefit of this multi-detector array imaging geometry isthat the depth of field of the object plane 280 is narrowly defined.Structures lying in front of or behind it will be blurred (out offocus). X-ray pencil beams from a first aperture 141 and a secondaperture 142 are depicted in FIG. 5 passing through an object plane 280a distance S₀ from apertures 141,142 and passing through a plane 281 adistance S₁ from apertures 141, 142 where S₁>S₀. The bubbles representimage pixels IP₁ through IP₀. As can be readily seen, the resolution atS₁ is less than that available at S₀. This feature provides for improvedlocalization and visualization of detailed structures in the plane ofinterest 280, while providing an adequate depth of field that may bemodified by the system geometry.

Conventional image intensifier technology typically has basicconstraints that limit a system's sensitivity. A scanning-beam x-rayimaging system can result in the subject under examination being exposedto the lowest possible level of x-rays commensurate with achieving imagequality adequate to meet the requirements of the procedure beingperformed. This means that the system used to detect the x-ray photonsemerging from the subject preferably has the highest possible detectivequantum efficiency. To achieve this, the scintillating material used inthe individual detector elements preferably has a length in thedirection in which the x-ray photons travel that is sufficient to ensurethat no x-ray photons emerge from the end opposite the incident x-rays,i.e., the x-ray photon energy should be adequately dissipated in thematerial to maximize the output of the detector.

There are several types of individual detector elements which can beused in the scanning-beam x-ray imaging system. Semiconductor detectorsusing silicon, selenium, cadmium telluride, cadmium zinc telluride, orother materials can be used. Scintillators or optical detectors can alsobe used, for example, cesium iodide or cadmium tungstate scintillatorswith amorphous-silicon, CMOS, or CCD optical detectors. A scintillatorin which x-ray photon energy is converted to visible light energy andthe light intensity is then converted to an electrical signal by meansof a photomultiplier, photo diode, CCD or similar device can beutilized. Because the information from each aperture must be obtained ina very short time period, the scintillating material should have a fastresponse and a minimum afterglow time. Afterglow is the phenomenonwherein the scintillator continues to emit light after the stimulatingincident x-rays have ceased. Even faster response and shorter afterglowtimes are required if x-ray intensity measurements are obtained usingthe x-ray photon counting technique.

One type of photon-counting detectors comprises detectors that convertincoming photons into charge carriers such as electrons or holes in asemiconductor material, or positive and negative ions in a gas or liquidmaterial. An electric field applied to the material will sweep positivecharges, such as positive ions or holes, towards one electrode andnegative charges, such as negative ions or electrons, towards anotherelectrode. As the charges accumulate on their respective electrodes,they form a current pulse. Such a current pulse can be analyzed usingpulse-height analysis techniques to yield a count of the number ofincoming photons.

With this type of photon counting detector, individual selected photonsthat strike the detector surface are counted by detecting a pulse ofelectrons or other charge carriers. The pulse of charge hassignificantly greater intensity than the charge detected betweenselected photons or when selected photons are absent.

A second type of photon-counting detectors comprises detectors thatconvert an incoming photon into several optical photons which togetherform an optical pulse. The optical photons can be in the visible rangewith wavelengths approximately within the range of from 400 nm to 700nm; infrared photons with wavelengths longer than approximately 700 nm;or ultraviolet photons with wavelengths shorter than 400 nm. The opticalpulse can be detected with an optical detector such as a CCD, aphoto-diode, or a photo-transistor, thereby transducing the opticalpulse into an electrical pulse. The electrical pulse can be analyzedusing pulse-height analysis techniques to yield a count of the number ofincoming photons.

The area of the circular active area of collimation grid 90 ispreferably larger than the area of multi-detector array 110. Thus theaxes of the x-ray pencil beams 100 emitted from the respective apertures140 of collimation grid 90 all converge toward the multi-detector array110 while each individual x-ray pencil beam 100 diverges, or spreads, aswould a flashlight beam to cover the face of the multi-detector array110.

Image reconstruction can be utilized to obtain high quality x-rayimages. The output of the multi-detector array is preferably not applieddirectly to the luminance input of a video monitor. Instead, digitizedintensity data for each image pixel are stored in a discrete address ina “frame store buffer”. More than one such buffer may be used in certainapplications. Pixel addresses within the buffer can be randomly accessedand the intensity value can be manipulated mathematically. This functionhas application in applying various image enhancement algorithms and itallows for pixel assignment of the data from discrete segments of thedetector array.

Referring to FIG. 6, this diagram illustrates the divergence of a singlex-ray pencil beam 100 from aperture 140 to the multi-detector array andhow it intersects an object 80 (not shown) at object plane 280. Imagepixel 121 can be just one of the image pixels comprising the x-raypencil beam intersection area 122 of object plane 280. A representativesample of the axes 102 of the x-ray micro-beams created by having asegmented array are also shown. In FIG. 6, x-ray pencil beam 100 isshown emitted through a single aperture 140 of collimator grid 90. X-raypencil beam 100 as it exits aperture 140 can diverge forming a conehaving a cross section the size of the aperture as it exits the apertureto a cross section covering the scintillators of the detector elementsof the multi-detector array by the time it reaches 96 elementmulti-detector array 110. 96 element multi-detector array 110 can bepositioned and designed such that the area of the cone of the x-ray beam100 just covers the surface area of the multi-detector array 160 whenthe x-ray pencil beam 100 intersects the face of the multi-detectorarray.

As x-ray pencil beam 100 passes through object 80, information aboutobject 80 can be detected by the multi-detector array 110 as x-rayintensity values. Because multi-detector array 110 is composed of 96separate detector elements, each detector element 160 can detect onlythe intensity value for the particular x-ray micro-beam 101 of a segmentof x-ray pencil beam 100 that it intersects with. The cross sectionalshape and area of the x-ray micro-beams can correspond to the crosssectional area and shape as the input face of the detector elements. Forexample, if the input faces are square, the x-ray micro-beam can have asquare cross section. The x-ray pencil beam 100 emitted from eachaperture 140 on collimator grid 90 can therefore generate one group of96 separate or discrete pieces of information (the intensity value ateach detector element) about 96 areas of object 80 in the x-ray pencilbeam's 100 path 122. The intensity information from each of the x-raymicro-beams can provide partial image pixel information which can beused to compile complete image pixel information for each image pixel ina desired plane of object 80.

FIG. 7 illustrates the axes 102 of all of the x-ray micro-beams from allof the apertures 140 that intersect a single image pixel 121 in objectplane 280 as they travel to the multi-detector array 110. This imagepixel group of x-ray micro-beams can be ultimately processed to generatean image pixel on a video monitor. In one embodiment of thescanning-beam x-ray system, the apertures 140 on collimator grid 90 cangenerate x-ray pencil beams 100 in a predetermined pattern. As x-raypencil beams 100 pass through an object, x-ray micro-beams 101 fromadjacent and nearby apertures can intersect at, for example, point 121(e.g. an image pixel) in the object. The intensity of each of thesex-ray micro-beams 101 from these x-ray pencil beams 100 after they passthrough the object can provide information about these intersectingpoints in the object. In this embodiment, each intersecting point on theobject can therefore be considered a single-image “pixel” 121. Inaccordance with the techniques explained in more detail herein, eachimage pixel 121 can be mathematically reconstructed from the intensityinformation of the separate x-ray micro-beams 101 that were generated bythe detector elements 160 for each of the emitted x-ray pencil beams 100from, for example, the image pixel group of apertures that generatedx-ray micro-beams whose axes passed through the object at that point,image pixel 121. In this example, a corresponding pattern of dataassignment is repeated as the scanning x-ray beam passes behind all ofthe pixels.

In the displayed image, with a sub-sampling ratio of 1:1, the numericalvalue of each image pixel is equal to the sum of “n” parts where “n” isthe number of detectors in the multi-detector array 110 (in thisexample, n=9). When constructed as shown in this example, themulti-detector array 110 together with the image reconstruction methodselected, has the effect of fixing the working distance at which optimumfocus is obtained and providing a plane of optimum focus.

The total area of the multi-detector array 110 should be large enough tointercept all of the x-rays in x-ray pencil beam 100 emanating from thecollimation grid 90, to avoid exposing the patient to x-ray radiationwhich does not contribute to the image. Outside of the plane of optimumspatial resolution, SO (280 in FIG. 5 and FIG. 6), spatial resolutionwill degrade. In some applications, degraded spatial resolution outsideof the depth of field of the system may be seen as being advantageousbecause blurring of detail outside of the area of interest may tend toincrease the perception of details within the area of interest.

A number of methods can be used to obtain a usable image from the dataobtained as described above. A simple convolution method may be used.Two additional methods can be utilized for obtaining maximal resolutionand sensitivity from the captured data, the multi-image convolutionmethod and the multi-output convolution method. An advantage of themulti-image convolution method over the multi-output convolution methodis that the former allows the plane of optimum focus to be selected insoftware after the data is captured while the latter does not. Thelatter method, however, may be performed quicker where timing is alimitation.

The scanning-beam imaging system described herein can be used togenerate a set of sequential planar images which can then be used toform a tomograph or a three dimensional display of the object 80. Animage set can be analyzed to produce a three dimensional imageconsisting of a series of images at various depths by re-analyzing thedata set with various values corresponding to planes of interest in theobject 80.

An alternative image reconstruction method can be employed toreconstruct images along multiple focal planes. This imagereconstruction method is referred to as m,n image reconstruction. Itwill be noted that there are numerous planes parallel to the sourceplane and detector plane where multiple beams pass throughregularly-spaced points in the plane. These planes are referred to asfocal planes or image planes. The regularly-spaced points are referredto as image pixels. Each focal or image plane comprise characteristicswhich differ from other focal planes, including distance from thesource, spacing of image pixels, and size of the image plane. Due topartial image reconstruction around the perimeter of the image, thenumber of fully reconstructed image pixels is slightly lower than theabove number and the total number of fully and partially reconstructedimage pixels is slightly higher than the above number. The m,n imagereconstruction method is more flexible than the previously describedreconstruction methods. As described, m,n image reconstruction cangenerate a wide variety of focal planes at numerous positions betweenthe source and detector planes. Many of the focal planes have a smallpitch between image pixels which can be used to produce images with highspatial resolution. The ability to reconstruct a wide variety of focalplanes can be used to move the focal plane with respect to the sourceand detector by simply selecting a suitable image plane near the regionof interest of the object to be imaged. The m,n image reconstructionmethod can also be used to increase the effective depth of field of animage by simultaneously reconstructing multiple focal planes around aregion of interest. The reconstructed planes can be combined to producea single image with high spatial resolution over a larger range ofdistances from the x-ray source plane. The multiple reconstructed planescan be combined, for example, by adding together only the high spatialfrequency components from each reconstructed plane.

Under one embodiment of the present invention, the imaging systemutilizes the sub-sampling method to process the detected information.The sub-sampling method can be employed in a reverse geometry scanningbeam x-ray system utilizing a sub-sampling ratio of 9:1 with amulti-detector array including ninety-six detector elements arranged ina pseudo-circle. The multi-image convolution method, the multi-outputconvolution method, m,n image reconstruction method and sub-samplingmethod is described more fully in U.S. Pat. No. 5,651,047 entitled“Maneuverable and Locateable Catheters” which has been incorporatedherein by reference in its entirety.

To generate an image pixel, the processed x-ray intensity valuesdetected by the multi-detector array 110 for each x-ray micro-beampassing through that image pixel IP are summed and output to a videomonitor. For image reconstruction using a sub-sampling ratio of 1:1 eachlogical detector element of the logical array is capable of providinginformation about each image pixel in the object. For imagereconstruction with a sub-sampling ratio of x:1, where x is a numbergreater than 1, less than all of the logical detector elements arecapable of contributing information about a particular image pixel. Theactual number capable of contributing information will depend on theparticular sub-sampling ratio selected. With a sub-sampling ratio of9:1, only 16 logical detector elements of the 144 logical detectorelement logical array will provide information about any particularimage pixel.

In the sub-sampling method with a sub-sampling ratio of 9:1, the logicalarray can include sixteen virtual detectors. The virtual detectors caneach include 9 logical detectors arranged in a 3 by 3 array.Alternatively, if a sub-sampling ratio of 4:1 were used, there would be36 virtual detectors, each including 4 logical detector elements. Usinga sub-sampling ratio of 1:1 there would be 144 virtual detectors eachincluding 1 logical detector element.

Each of the 16 logical detector elements used to reconstruct a singleimage pixel using a sub-sampling ratio of 9:1 can be in differentvirtual detectors. Each virtual detector contributes partial image pixelinformation for nine different image pixels. Complete image pixelinformation is obtained by combining the information from the logicaldetectors in the same virtual array location from all 16 virtualdetectors.

Additional image reconstruction methods and techniques can be utilizedto generate information for a wide variety of planes and slices atnumerous positions between the source and detector. These methods andtechniques are described more fully in U.S. Pat. Nos. 6,178,223 and6,181,764, entitled “Image Reconstruction Method and Apparatus” and“Image Reconstruction for Wide Depth of Field Images,” all of which arehereby incorporated herein by reference in their entirety.

The images can be acquired, including reconstruction, and the exposurerates can be optimized using the methods described above and below,rapidly enough to create a continuous, real time video representation ofthe motion of the object, including organs such as one or more of thelungs, the heart, or other organs, or instruments, such as catheters orstents, or implantable objects such as valves, in real time.

The scanning-beam imaging system described herein can be used togenerate a set of sequential planar images which can then be used toform a tomograph or a three dimensional display of the object 80. Animage set can be analyzed to produce a three dimensional imageconsisting of a series of images at various depths by re-analyzing thedata set with various values corresponding to planes of interest in theobject 80.

FIG. 8 is a diagram illustrating one embodiment of the presentinvention. In FIG. 8 electron beams 1301 a and 1301 b impinge on focalpoints 1303 a and 1303 b of target 1307, and electron beams 1302 a and1302 b impinge on focal points 1304 a and 1304 b of target 1307. Here,for simplicity, only one row of focal spots is shown, whereas in apreferred embodiment the target may be made of several rows of focalspots. When the electron beams 1302 a and 1302 b strikes target 1307 atfocal spots 1304 a and 1304 b, x-ray beams 1309 a and 1309 b are createdand measured by detector(s) 1308. An object, e.g. a human patient, to beimaged contains some region of interest 1306. Region of interest 1306 isshown here for simplicity as a single region, but in actuality could be2, 3, 4, or more distinct regions of interest. In order to minimize thedose to the patient and reduce exposure to medical personnel, only thefocal spots 1304 a and 1304 b and those between them are illuminated bythe full electron beams 1302 a and 1302 b. As focal spots 1303 a and1303 b would expose an area outside of region of interest 1306, patientdose is reduced by reducing electron beams 1301 a and 1301 b orredirecting the beams away from focal spots 1303 a and 1303 b. Thisprocess is referred to as region of interest filtering. It can be morespecifically described as digital or electronic region of interestfiltering, as the definition of at least one region of interest 1306does not require the placement of mechanical components such asshutters, but instead is implemented by electronically controllingelectron beams 1301 a, 1301 b, 1302 a and 1302 b.

In an embodiment of the present invention, detector 1308 is an energyresolving detector with two or more energy bins, preferably 10 or lessbins, more preferably 5 or less bins, and most preferably 2 bins, andthe contrast-to-noise ratio is optimized for a given contrast medium byweighting the number of detected x-rays in each bin by using theexpression

${w(E)} = \frac{1 - e^{{- {\lbrack{{\mu_{c}{(E)}} - {\mu_{b}{(E)}}}\rbrack}} \cdot d}}{1 + e^{{- {\lbrack{{\mu_{c}{(E)}} - {\mu_{b}{(E)}}}\rbrack}} \cdot d}}$as the weighting factor, wherein d is the thickness of the contrastmedium, μ_(c)(E) is the energy dependent attenuation coefficient of thecontrast medium, and μ_(b)(E) the attenuation coefficient of thebackground.

In another embodiment of the present invention, the exposure to x-raysof at least one person is optimized by modifying target 1307, andthereby modifying the shape of the x-ray energy spectrum, to best matchthe x-ray energy spectrum to one or more of: the needs of the procedureto be performed; the properties of the subject being imaged; theproperties of the target organ to be imaged; any instrument that will beused during the procedure; and any contrast agents used during theprocedure. The modification of target 1307 can involve the use ofmaterials, wherein those materials are chosen from materials includingbut not limited to tungsten, copper, aluminum, beryllium, lead, rareearth elements including but not limited to gadolinium, and alloys ormixtures thereof.

It is an aspect of an embodiment of target 1307 that target 1307comprises a tungsten layer in contact with the vacuum layer into whichelectron beam 1302 a and 1302 b impinges and creates x-rays 1309, andwherein the x-ray spectrum is modified by one or more layers of anadditional material or materials. A modified target 1307 can comprise aberyllium sheet onto which a film of tungsten is deposited, whileberyllium is in contact with a thin layer of cooling water, which isfollowed by thin aluminum sheet, and wherein the procedure may utilizeiodine as a contrast agent, and wherein the procedure may be a cardiacintervention. FIG. 9 is a table displaying contrast-to-noise improvementfor different iodine concentrations. In a further embodiment, there isat least one additional layer comprising at least one rare earth metalbetween the tungsten film and the beryllium sheet, where the at leastone rare earth metal can comprise gadolinium.

There are many ways that region of interest 1306 can be defined. In apreferred embodiment, a healthcare provider is presented with an image(a recently acquired frame) from which to define region of interest 1306or multiple regions of interest. The image from which the health careprovider defines region of interest 1306 and other regions of interestthat will be exposed to x-rays can be the entire field of view of thesystem, or any subset thereof, including a previously defined region ofinterest. The health care provider can define the region(s) of interestin many ways, but in a preferred embodiment the region(s) of interest(s)are defined by the health care provider drawing the region on a displaydevice, preferably using a dedicated stylus but in other embodimentsusing a finger or any other object.

In another embodiment of the present invention, the control and imagingportions of the apparatus are used to automatically define a region ofinterest that is optimal for the procedure to be performed, usingmethods chosen from a list including but not limited to: identifying anorgan of interest, identifying a tumor, utilizing image data takenpreviously of the patient, using an algorithm that uses patient dataincluding but not limited to height, weight, body mass index, age, sex,race, medical status or conditions, chest size, and/or length of arms orlegs.

To further reduce the x-ray dose delivered to the patient as well asreduce the exposure to x-rays of health care providers, one embodimentof the present invention determines the minimum amount of x-rayradiation required to make a reconstructed image with a sufficientsignal-to-noise ratio and resolution for the procedure to be performed,in at least one region of the object, preferably in more than oneregion, e.g. in region 1311 a and region 1311 b. Further, the apparatusthen limits the x-ray dose to regions 1311 a or 1311 b to that minimumrequired amount. This is referred to herein as equalization filtering.

As shown in FIG. 8, there are two separate regions 1311 a and 1311 b towhich equalization filtering may be applied. To apply equalizationfiltration, the x-ray absorbance of regions 1311 a and 1311 b aredetermined by exciting focal spots 1304 a and 1304 b with electron beamsof known qualities. This causes the formation of x-ray beams 1309 a and1309 b which pass through regions 1311 a and 1311 b and are subsequentlydetected by detector 1308. Based on this measurement, electron beams1302 a and 1302 b are modified in a way that reduces the total x-raydose to regions 1311 a and 1311 b to only that needed to achieve thesignal-to-noise ratio required at the image resolution necessary tocomplete procedure being performed. The adjustable properties ofelectron beams 1302 a and 1302 b may be chosen from a list including butnot limited to voltage, current, or duration. In a preferred embodimentthe electron beams are raster-scanned over the focal spots on target1307, dwelling on each targeted focal spot for a predetermined amount oftime. The reconstructed image is formed from 2 or more raster scans ofthe electron beams over target 1307, and the equalization filtration isimplemented by adjusting the number of times the electron beams dwell oneach focal spot 1304 a and 1304 b. This method is herein referred to asdigital equalization filtration.

In one embodiment of the present invention, the dose of x-rays to apatient is minimized by determining the amount of x-ray radiationdetected by one or more pixels in one or more detectors, e.g. detector1308, during an image acquisition and terminating the delivery ofradiation to a region of the patient, e.g. region 1311 a or 1311 b,based on a threshold amount of measured x-ray radiation detected by theone or more detectors or detector pixels. In the digital equalizationfiltration method described above, this is implemented by measuring theamount of x-rays being detected by at least one detector or at least onepixel in at least one detector and during dwells of electron beam 1302 aand 1302 b on a focal spot 1304 a and 1304 b; determining a sum total ofx-rays detected during one or more dwells of electron beam 1302 a and1304 b on a focal spot 1304 a and 1304 b; and skipping future dwells tobe performed on that focal spot 1304 a and 1304 b during that imageacquisition when the sum total exceeds a predetermined threshold amount.This process is repeated for each focal spot 1304 a and 1304 b to beused. This embodiment has the advantage of a high signal-to-noise ratiofor the determination of whether the signal measured has exceeded thethreshold value.

In another embodiment of the present invention, the amount of x-raysdetected during the first dwell of x-ray beam 1302 a and 1304 b on eachfocal point 1304 a and 1304 b is used to then compute the expectednumber of dwells required to measure a sufficient amount of x-rays tocreate an image of the required signal-to-noise and resolution, and thatnumber of dwells is used during the imaging process. This determinationis repeated for each focal spot to be used. This method has theadvantage of being less computationally intensive but suffers from lowersignal-to-noise than the previous embodiment.

Due to software or hardware limitations, it may be preferable to limitthe number of changes in the number of dwells between subsequent focalspots during the raster scan of target 1307. This can be achieved by analgorithm that allows a range of values for the threshold value, forexample by putting in some hysteresis that stops the number of dwellsfrom bouncing up and down due to noise, for example Poisson or “shot”noise, in the number of x-rays arriving during the dwells during theraster scan.

It is preferable to perform equalization filtration in real time, i.e.during each image acquisition, by adjusting the exposure time or someother property of the electron beam to optimize the number of x-raysemitted from each focal spot as the imaging system runs. This has theadvantage of not only minimizing the x-ray dose to the patient but alsoproviding an optimized reconstructed image for each acquisition, forexample each frame in a video image, by removing artifacts and issuesdue to, for example, heart rate and/or respiratory rate. However, thepreviously discussed embodiment of the present invention that minimizesthe dose of x-rays to a patient by determining the amount of x-rayradiation detected by one or more detectors or pixels in one or morepixelated detectors during an initial image acquisition, and based onthat measurement, delivering an optimal, reduced, amount of x-rayradiation to a region, e.g. region 1311 a or 1311 b, of the patientduring one or more subsequent image acquisitions has the advantage ofbeing less computationally intensive.

Other data and methods can be used to implement and refine equalizationfiltration, using methods chosen from a list including but not limitedto: identifying an organ of interest, identifying a tumor, utilizingimage data taken previously of the patient, using an algorithm that usesa predefined model of a region to be imaged and optionally includespatient data including but not limited to height, weight, body massindex, body fat percentage, age, sex, race, medical status orconditions, and/or size of a body region to be imaged, including but notlimited to thickness, length, height, volume, and/or circumference,including but not limited to the size of the subjects chest, abdomen,waist, neck, head, arms, or legs, and using this information to modifythe amount of radiation delivered to the various regions to be imaged.These techniques are preferably combined with the digital equalizationfiltration embodiments described above and/or combined with theelectronic region of interest filtering described above.

In one embodiment of the present invention, image equalization isimplemented using information about a reconstructed image, said imagebeing a previously acquired image, e.g. an earlier frame. In thisembodiment, it is preferable that the previous image be a recentlyacquired image, ideally in the range of 1 to 20 images previouslyacquired, and may be chosen from 1, 2, 3, 4, 5, 10, 15, or 20 imagesprevious. It may be advantageous to utilize only image data from asubset, or region, within the final reconstructed image.

In another embodiment of the present invention, equalization filtrationis implemented using a preliminary, interim reconstruction generated byx-rays detected from one or more detectors or one or more pixels in oneor more detectors during a subset of the final dwell time on one or morefocal spots. It is preferable that the subset of the final dwell timeincludes the first dwell and may optionally include data from the firstand second; first, second, and third; first, second, third, or fourth;or any other combinations of the available dwells, including but notlimited to during the first 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15,16, 17, 18, 19, or 20 dwells, or during the first ¼, ⅓, ½, ⅔, or ¾ ofthe maximum number of dwells.

For digital equalization filtration, it is preferred that every time theelectron beam dwells on focal spot 1304 a or 1304 b, the length of timeit dwells on focal spot 1304 a or 1304 b is characterized by apredefined dwell time chosen from zero milliseconds (if region ofinterest filtering is simultaneously used and has determined that focalspot 1304 a or 1304 b does not illuminate the region(s) of interest) toa predetermined non-zero dwell duration. The total dwell time for focalspot 1304 a or 1304 b may then be the predetermined non-zero amount oftime multiplied by the number of non-zero dwell time dwells for focalspot 1304 a or 1304 b, and the dwell time for other focal spots may besimilarly determined. The amount of delivered x-ray radiation can beadjusted by adjusting the number of non-zero duration dwells. Theproperties of the electron beam including but not limited to voltage,current, focusing, or dwell time per dwell can be the same for all focalspots used in target 1307, but these may also be adjusted to implementor refine equalization filtration.

For the digital equalization filtering described herein, it is preferredto reconstruct each image based on both the amount of x-ray radiationdetected by the at least one pixel of the at least one detector 1308,and the number of non-zero dwell time dwells for each correspondingfocal spot 1304 a and 1304 b. In an embodiment of the present invention,the amount of radiation detected by the at least one detector 1308summed over the total number of dwells per image is divided by thenumber of dwells or the total dwell time of the electron beam 1302 a or1302 b on the corresponding focal spot 1304 a or 1304 b and thisnormalized detected x-rays per dwell or per unit time is used insubsequent image reconstruction.

X-ray detector 1308 can include any usable technology, and may be chosenfrom a list including but not limited to: a photon counting detector, acharge-integrating detector, or an energy resolving detector.

In an embodiment of the present invention, detector 1308 is a photoncounting detector, and the dwell time and/or the number of fixed dwellsis modified by truncating the exposure of focal spot 1304 a and 1304 bto electron beam 1302 a and 1304 b when the number of photons detectedby detector 1308 exceeds a pre-determined threshold number, andsimilarly truncated for other focal spots.

In an embodiment of the present invention, detector 1308 and/or otherdetectors are photon counting detectors or more preferably energyresolving detectors, and the dwell time and/or the number of fixeddwells is modified by truncating the exposure of focal spot 1304 a and1304 b to electron beam 1302 a and 1304 b when the number of photonsdetected by detector(s) 1308 exceeds a pre-determined threshold numberwherein the threshold number is modified based on one or both of: thenumber of photons detected in at least one energy bin by at least onepixel or detector 1308 during a previously dwelled upon different focalspot or focal spots; or the number of dwells before the threshold wasreached during exposure of a previously dwelled upon different focalspot or focal spots during a given raster scan. This has the advantagethat the threshold can be modified to minimize the number of changes inthe number of dwells between subsequent focal spots in the scan oftarget 1307, simplifying the mechanism and the method, and/or allowingit to run faster.

Under an alternative embodiment of the present invention, x-ray dosageand radiation on object 80 or the patient and attending staff can befurther reduced by use of adaptive methods and apparatus. The imagingsystem may or may not be used in conjunction with a radiation source andmay or may not be part of a radiation therapy system. X-ray imaging ofthe human body involves x-rays penetrating through different regionswith highly varying attenuations; the intensity of x-rays reaching adetector depends on the amount the x-rays from a source were scatteredand absorbed within the patient volume. Areas of the x-ray image can beoverexposed with the result that object 80, the patient and attendingstaff can be exposed to unnecessary dose or x-ray radiation.

The adaptive methods and apparatus of one embodiment of the presentinvention can reduce x-ray dose and radiation and can be utilized forreal-time x-ray imaging. Rather than acquiring the full-field image witha single exposure, the imaging system can use many exposures orprojections of small areas of object 80 to generate the image, thenumber of exposures or projections can be as high as thousands. For a7-inch field of view, up to 9,000 exposures or projections can be used.Multiple exposures or projections are possible through use of ascanning-beam x-ray source or a multi-pixel carbon nanotube x-ray sourceor discrete cathode x-ray source or other multi-pixel x-ray sources assource 10. Each exposure or projection can significantly overlap withneighboring exposures or projections, thus the exposures or projectionscan be redundant. Rather than using the same exposure for everyprojection, the imaging system exposes every projection with justsufficient exposure to obtain the desired image quality. For example,lungs may be a body part requiring less exposure to obtain an image of agiven quality, as their attenuation characteristics permit thepenetration of a relatively large percentage of incident x-rayradiation. Therefore, in an embodiment of the present inventionutilizing a scanning-beam x-ray source or a multi-pixel carbon nanotubex-ray source or discrete cathode x-ray source or other multi-pixel x-raysources, the imaging system may expose the lung to less radiation thanother body parts that are relatively x-ray opaque, such as the heart,and still achieve sufficient image quality. The imaging system canadjust exposures and radiation by using different tube voltages, beamcurrents, exposure times, repetitions of fixed-length exposures, or thecombination of the former. One setting can be 120 kV tube voltage at 17kW power and another setting can be 80 kV tube voltage at 9 kW power.Reducing the exposure or radiation is not the only benefit. The imagingsystem also can define preselected regions of interest for increasedimage quality and other regions with lower image quality. Thus, regionsof interest that require high image quality, can receive more exposurethrough increased repetition of fixed-length exposures of that region,different tube voltages, beam currents, exposure times, or thecombination of the former.

Under an alternative embodiment of the present invention, source 10 canbe a scanning-beam X-ray source. Electron beam 40 can be swept acrossX-ray target 50 and dwells at distinct focal spot 60 positions for apredetermined exposure time, which depends on the frame rate and thesize of field of view. At a frame rate of 15 fps (7″ mode field ofview), the exposure time is 8 μsec. This exposure time is broken into 1μsec illuminations (rescans) separated in time. For given materialselection and dimensions of x-ray target 50, use of 1 μsec illuminationscan keep the temperature of x-ray target 50 below desired maximums(maximums being set by the potential of the target to “burn out” whenoverheated by the incident beam in a localized region). Thus, everyfocal-spot position is rescanned several times. Different size or aspectratio of field of view can also be used, 5″×5″, 5″×10″, etc, or, 6″, 7″,8″, 9″, 10″, 11″, 12″, 13″, 14″ or 15″.

The imaging system can utilize equalization or region of interestfiltration or both. For both equalization or region of interestfiltration, offline implementation or real-time implementation can beutilized. The imaging system can acquire images from up to 9,000different focal-spot positions for each full-field image. Eachacquisition can be repeated up to 8 times (8 rescans). With 7″ field ofview, this results in a frame rate of 15 fps. Alternatively, eachacquisition can be repeated 2, 3, 4, 5, 6, 7, 9, 10, 11, 12, 13, 14, 15,16, 17, 18, 19 or 20 times (rescans). The imaging system can aggregatethe rescans or it can save each rescan separately (rescan images). Eachrescan can be saved as a separate file and can be buffered.Alternatively, the rescans can be aggregated immediately.

Under an alternative embodiment of the present invention, the imagingsystem can implement equalization filtration using separate oraggregated rescan images. In this embodiment, the imaging system mayfirst determine a target number of photons per detector image. It maythen determine the number of photons in the rescan image(s) andaggregate rescan images until at least the target number of photons isreached or all 8 rescans are added. This procedure yields a “rescan map”detailing how many rescans are needed per focal spot position and amodified detector image file that can be reconstructed with thereconstruction engine.

Equalization filtration not only saves dose but can be a very effectiveway to compress dynamic range and thereby improve image quality. Aspreviously discussed, equalization filtration can dynamically andautomatically vary the exposure depending on the opacity of the regionexposed. Hence, it can compress the dynamic range by reducing exposuresignificantly in translucent areas such as the lung field and maintainexposure in more opaque regions.

Under another embodiment equalization is performed by variation of thebeam current and the beam current is adjusted according to the fluxmeasured at the detector. Importantly the beam current can start at lowvalues to not saturate the detector even with no attenuator, e.g. nopatient, present. With an attenuator present this low beam current willresult in a very low flux being detected at the detector. Feedbackrelaying detected flux can lead to up-regulation, i.e. increasing, ofthe beam current until desired flux is achieved. The up-regulation canbe done during a single exposure or implemented via rescans.

FIG. 10 is a plot illustrating a manner of equalization filtration ofone embodiment of the present invention. In this embodiment,equalization filtration can be implemented by performing the first scanacross the entire field of view or part of the field of view. A “mainthreshold” can be determined to limit exposure in each partial image ofeach rescan. Once desired exposure level or threshold has been achieved,the x-ray beam in partial images can be turned off or the x-ray beamsassociated with particular collimator holes can be turned off. Withproper threshold value, partial images or collimator holes receivingexposure above the threshold will be turned off. In this manner, thedynamic range of exposure values can be compressed and dose can besaved. Image quality can be improved by reducing exposure in the lightareas thus reducing the dynamic range.

In FIG. 10, the x-axis, labeled “Collimator Hole” represents theposition of the scanning beam over subsequent apertures in collimationgrid 90; it could also be viewed as a temporal axis, where the units oftime are equal to the dwell time for a single aperture. The y-axisindicates the amount of x-ray flux emitted from an aperture that reachesdetector 110 after penetrating some amount of patient volume. It can beseen that aperture 140 will be scanned by beam 40 or otherwiseilluminated, i.e. “on,” until the amount of x-ray flux from a givenaperture reaching detector 110 exceeds the main threshold, at whichpoint the electron beam is switched “off.” If the flux drops below themain threshold, the electron beam is switched on again, and so forth.

Under an alternative embodiment of the present invention, equalizationfiltration can be done by calculating the rescan map in real time whilethe image is acquired. Implementation of the rescan map allows theimaging system to perform an equalization filtration scan in hardware,turning on or off the x-ray beam depending on how many rescans areneeded. A complication is that there may be a hardware limit related tothe grid control on the electron beam that limits the number of timesthat the x-ray beam can be turned on or off (number of switches) duringeach image acquisition frame. Such a limitation results from impedance,resistance, capacitance and inductance characteristics of the hardware.To limit the number of switches, equalization filtration can be modifiedby not only relying on a single threshold but rather considering a bandof target counts, effectively low-pass filtering the rescan map. With asingle threshold as in FIG. 10, the imaging system may switch 774 timesor more. With a threshold band of target counts, the imaging systemcould reduce the number of switches to 548 or some other number below774 or the number of switches experienced with a single threshold. Withadjustment of threshold settings in the threshold band, the imagingsystem could further reduce the number of switches to 463 or some othernumber below 774 or the number of switches experienced with a singlethreshold.

The imaging system can implement equalization filtration using athreshold band, comprised of an upper maximum threshold and a lowerminimum threshold, of target counts by saving or tracking exposure(flux) for each area or collimator hole in the rescan map. As theimaging system performs a scan, when the exposure or flux for an area orcollimator hole increases from a level below a maximum threshold to alevel just above the maximum threshold, the x-ray beam for the area orcollimator hole is turned off. When the exposure or flux for an area orcollimator hole falls below the maximum threshold, the x-ray beam forthe area or collimator hole remains off. When the exposure or flux foran area or collimator hole decreases further to a level below a minimumthreshold, the x-ray beam for the area or collimator hole is turned on.When the exposure or flux for an area or collimator hole increases abovethe minimum threshold, the x-ray beam for the area or collimator holeremains on. When the exposure or flux for an area or collimator holeincreases above the maximum threshold, the x-ray beam for the area orcollimator hole will be turned off. At different angles, the imagingsystem can save between 33% and 60% dose with an average dose saving of47% using equalization filtration. Alternatively, dose saving of 41%,43% or 45% of primary photons can be achieved.

If visualized using FIG. 10, the threshold band may appear as ahorizontal line positioned higher on the y-axis than the main thresholdrepresenting the maximum threshold and a horizontal line positionedlower than the main threshold representing the minimum threshold. Thestate of the collimator hole would only switch from “on” to “off” whenthe flux transitioned from below to above the maximum threshold andwould only switch from “off” to “on” when the flux transitioned fromabove to below the minimum threshold.

Under an alternative embodiment of the present invention, equalizationfiltration can be done by calculating the rescan map from a single frameof a real-time video and use it for the subsequent frame. Implementationof the rescan map in this manner allows the imaging system to perform anequalization filtration scan in hardware, thus turning on or off thex-ray beam depending on how many rescans are needed. This embodiment canbe used with a single threshold or a threshold band.

Under an alternative embodiment of the present invention, the imagingsystem can utilize region of interest filtration, and two or moreregions of interest can be selected. The first region of interest maymaintain full image quality, with 8 rescans or the maximum number ofrescans, and the second region of interest may only be imaged atsignificantly reduced image quality, with 1 or 2 rescans or a lowernumber of rescans than full image quality. Outside the second region ofinterest, the beam may turned off entirely via electronic collimation oranother method. The rescan images can be aggregated based on theselection by the user or based on automatic or predetermined selectionby the imaging system. The resulting detector images can bereconstructed and can be evaluated for image quality and dose savings.The user can toggle between region of interest filtration and standardacquisition mode on a frame by frame basis. Collimation can be but doesnot have to be toggled as well.

Region or regions of interest can be selected by the user or selected bythe imaging system based on automatic or predetermined selection. Afterselection, the imaging system can define and implement region or regionsof interest by deploying a mechanical shield or shutter to blockelectron beam 40 before x-ray target 50 or the x-ray beam after x-raytarget 50. Shield or shutter can be made from tungsten, tungsten copper,tungsten alloy, lead, lead antimony, lead alloy, tantalum, tantalumalloy or other material with a high atomic number. (Materials with ahigh atomic number may successfully shield x-rays more so than materialswith low atomic numbers.)

Under an alternative embodiment of the present invention, region orregions of interest can be selected by the user or selected by theimaging system based on automatic or predetermined selection from avariety of shapes including without limitation, circle, oval, ellipse,square, rectangle, triangle, polygon and quadrilateral shapes. Theregion or regions of interest can also be the entire field of viewexcept the regions within the shape selected or the regions excluded bythe shape. The size of the region or regions of interest can be any sizeor range of sizes up to the size of the field of view. In one embodimentof the present invention, the region of interest is set to the size of ahuman heart or one chamber of the human heart and is intended to exposeonly the heart or selected chamber to radiation. In another embodimentof the present invention, the region of interest is the size of humanovaries and is intended to exclude radiation from the ovaries.

To implement region of interest filtering, the imaging system canperform an initial scan across the entire field of view. Alternatively,the imaging system can utilize a previous saved scan or image. Theimaging system can also perform one or more rescans across the entirefield of view. The imaging system can then perform one or more rescansbased on the shape selected or shape excluded at the size of the regionof the interest. The rescan would only involve exposure for some area orcollimator holes corresponding to the shape selected at the size of theregion of interest. The rescan may involve all of the collimator holeswithin the shape selected or a portion of those collimator holes butwould not involve areas or collimator holes outside the shape selected.In the example of a shape excluded, the rescan would only involveexposure for the area or collimator holes outside the shape excluded atthe size of the region of interest. The imaging system can also performone or more rescans based on one or more additional shapes selected orshapes excluded at one or more sizes of the region of the interest.

Under one embodiment of the present invention, the imaging systemperforms one initial scan and one rescan across the entire field ofview, two rescans or one additional rescan across a rectangular regionof interest corresponding to half of a human heart, and eight rescans orsix additional rescans across a smaller region of interest within therectangular region of interest. The entire field of view receives oneinitial scan and one rescan. The rectangular region of interest receivesone initial scan and two rescans and the smaller region of interestwithin the rectangular region of interest receives one initial scan andeight rescans. Dose savings of up to 82% of the primary photons can beachieved utilizing this region of interest approach.

Under an alternative embodiment of the present invention, the specificregion of interest can vary in position in the field of view over time.The region of interest can be a selected shape (e.g. circle, oval orellipse) of a specified size that follows a contrast agent as it flowsthrough vessels in the heart. The region of interest can beautomatically selected by, for example, following the tip of a catheteror by following the bolus of a contrast medium. The region of interestcan follow the edge of the contrast agent as it moves through vessels inthe heart. The region of interest can also be time-based in a mannerwhich tracks the motion of an organ such as the heart. The region ofinterest can be the area of the field of view experiencing motion.

Alternatively, the region of interest can be calculated by the imagingsystem. In a tomographic or tomosynthetic application, the user canselect a single plane or multiple planes of interest. The user canselect the region or regions of interest in the selected plane orplanes; the region of interest can be selected by the user in threedimensions. The imaging system can then track the region of interest orregions of interest in following frames by incorporating motion withinthe field of view and calculating the relative position of the region ofinterest or regions of interest. Under an alternative embodiment of thepresent invention, equalization filtration can be combined with regionsof interest to achieve maximal dose savings without detracting from thehigh image quality desired.

If target 50 includes a tungsten film, x-ray may be emitted through twomechanisms: Bremsstrahlung radiation, a direct results of the suddendeceleration of electrons by collision with heavier target particles,and K-line radiation, a result of incident electrons knocking targetelectrons out of their atomic “K” shell and other target electronsshedding energy to fill these created vacancies. Bremsstrahlungradiation takes on a spectrum of wavelengths and is the primarymechanism by which x-rays are emitted whereas K-line radiation emits aspecific wavelength characteristic to the atomic structure of tungstenor target material and is the less prevalent mechanism.

Under an alternative embodiment of the present invention, the beamhardening filter is brought into close proximity of the tungsten film inx-ray target 50 and emitted K-line radiation can have a similar spatialextent as the primary (Bremsstrahlung) radiation and can be useful forimaging. X-ray target 50 can also have an additional film with a rareearth that has K-lines particularly well suited for imaging.Specifically, the rare earth can provide large contrast in iodine, acontrast medium used in interventional cardiology. Various rare earthscan be used (e.g. Er, Gd, Dy, Sm) with different thicknesses (10 μm to150 μm) and X-ray tube potentials (60 kVp to 90 kVp). A 150 μm layer oferbium in x-ray target 50 at 80 kVp can result in a low dose. The dosecan be 87% of that found with 1 mm iodine using 0.1 mm Copper beamhardening filter, 70 kVp X-ray tube potential and maximum electron beamcurrent of 210 mA. A reduced Er layer thickness of 120 μm can also beused and Er can be used in x-ray target 50. For the beam hardeningfilter the dose can be 96% with 1 mm iodine using 0.1 mm Copper beamhardening filter, 70 kVp X-ray tube potential and maximum electron beamcurrent of 210 mA and for the integrated case it can 89%. Thus, using Erin the target can save 7.5% of the dose compared to using it as a beamhardening filter.

From a dose standpoint, a 150 μm Erbium target can be utilized. Er canbe a significant thermal resistor that slowed the time to equilibriumtemperatures significantly. A more convenient measure of time toequilibrium is the time to maximum temperature in the Be layer. With 1mm iodine using 0.1 mm Copper beam hardening filter, 70 kVp X-ray tubepotential and maximum electron beam current of 210 mA, this time can be25 μs and for a 20 μm layer of Er the time can almost double. At thetarget thickness of 150 μm, the time can increase to 520 μs. Long timesto equilibrate can lead eventually to overheating of x-ray target 50.

Real-time computing problems often require multiple graphics processingunits (GPUs) to provide sufficient computing power to performcalculations in an allowable time budget. Incoming data is oftenstreamed to these GPU devices at a continuous rate leading to severeperformance losses on the GPU. These performance losses result from thefact that a single GPU must split resources between theaggregation/pre-processing of the data set and the subsequentcomputational processing of that complete data set. To overcome thisproblem, one GPU 1104 can be used as pre-processing engine, and a set of1 to N (with N being a number between 1 and 100) GPUs as computationalprocessing engines, e.g. reconstruction GPU 1105. The pre-processingengine captures the continuous data stream and aggregates itappropriately. This aggregated and pre-processed data is thendistributed as a complete package to the processing GPUs using a singleburst memory operation. Rather than receiving the data as a continuousstream, these processing GPUs receive a data package in short timeinterval only briefly interrupting its internal computational work. Thisallows the processing GPUs to almost perform at 100% of their ratedperformance by fully occupying the available computational resources fora longer time period.

FIG. 11 is a diagram showing an overall architecture of one embodimentof the present invention. Stage I GPU 1104 accumulates and pre-processesdata directly from multiple SDM sources, then copies over the completeset via PCI-E bus 1115 to Stage II image reconstruction GPU(s) 1105 forthe computational work. In one embodiment, pre-processing GPU 1104 canbe a professional GPU with remote direct memory access (RDMA1)capability and the set of processing engines 1105 can be consumer gradeGPUs without RDMA. In this case, the memory transfer betweenpre-processing GPU 1104 and processing GPUs 1105 is orchestrated by thehost CPU. In another embodiment, all GPUs are professional-grade GPUsenabling memory transfers without host involvement. This has theadvantage that the process is independent of the timing imposed by theoperating system.

In one embodiment of the present invention, an array of GPUs toreconstruct data from an X-ray detector coming at a rate of 1 MHz. Eachdetector image contains of 160×80 bytes. Pre-processing GPU 1104receives the detector data via RDMA1 and aggregates the detector data toframes of 100×100×160×80 bytes for 10 inch×10 inch array of apertures140. This complete frame data is send to the processing GPUs 1105 at arate of up to 30 Hz.

In one embodiment of the present invention, the architecture is used ina low-dose fluoroscopic system that has a large field of view (FOV), asrequired for use in interventional radiology. The system is expected toreduce entrance X-ray exposure and the radiation dose to the patient bya factor of 4 and the occupational dose by at least a factor of 2 atequal image quality as compared to a conventional system. The system isbased on the Scanning-Beam Digital X-ray (SBDX) fluoroscope, a cardiacfluoroscopy system. Conventional fluoroscopy systems use an X-ray tubewith a single focal spot and an extended detector, with the patientpositioned close to the detector. In contrast, the imaging system canutilize an extended X-ray source that produces multiple focal spots. Theradiation is focused from every focal spot position onto a pixelateddetector with a custom collimator. The patient is positioned relativelyfar from the detector. Importantly, the final image is not generatedfrom a single detector image as in conventional systems. Rather, it iscomprised of up to 9,000 or 10,000 detector images that arereconstructed in real time. The largest gain in FOV is possible byemploying two laterally spaced detectors compared to one detector in theSBDX cardiology system. A collimator can be used to illuminate bothdetectors by alternating the illumination pattern on a row by row basis.

In one embodiment of the present invention, one imaging frame every 130ms can be produced. An imaging frame comprises 100×100 detector imagesof 160×80 detector pixels. The data are received by stage I GPU 1104continuously and stored in an llh data buffer or accumulation databuffer that is transferred to stage II GPU 1105 once a frame iscomplete. The GPU reconstruction kernel uses as inputs the aggregatedllh data buffer or accumulation data buffer, 2D detector data sets foreach detector, the detector hole sum data, the current z focus value,and the amount to offset each output data image in the finalreconstructed output.

GPU 1105 organizes its computation in blocks of worker threads, each ofwhich can perform its own independent calculations and memoryoperations. These threads can operate on four basic types of memory;global, shared, texture, and local (register). The fastest memory iskept in thread local registers, and is a very limited resource onlyvisible by the current working thread. Texture memory is read-only forthe life of the workload, and can be accessed by all threads in allblocks with very little latency. Unfortunately this texture memory islimited and the slightly slower shared memory is only available to allthe threads in a distinct thread block. Because most of the input dataresides in the slowest global memory, a strategy to read/write from thisspace in non-random patterns must be implemented in order to completethe process in time. One such strategy is to load values in a vectorizedfashion, which means a larger chunk of contiguous global memory isloaded, then broken down into subsets in thread local register memory.

The current llh input data buffer or accumulation data buffer can bestored in a format which splits the rows of data, mapping the rightdetector data to all odd rows, and the left detector data to all evenrows. Each thread block uses this odd/even information to determinewhere in the output buffer the weighted values are to be added. Eachthread in the GPU reconstruction kernel loads its required data incoalesced vectorized 4-element sections from global memory. Values whichwill be re-used by all threads in a particular block are stored locallyin shared memory. The input values which are specific only to thecurrent working thread will be loaded from global memory by theefficient vectorized method and stored in local register memory.Performance can be increased by allowing the kernel to load 16 llh framevalues or accumulation frame values in a single operation by casting theload as the 16 byte unit 4 type. Each worker thread in a block loads atotal of 32 values from the llh buffer or accumulation buffer byexecuting 2 contiguous 16 bytes loads. When writing to the outputbuffer, the reconstruction kernel makes use of the improved performanceof the atomic operations on global memory. During this process eachthread performs as many as 256 writes to global memory via theatomicAdd( ) operation. This method prevents a ‘race’ condition whicharises when more than one working thread attempts to update the samememory location at the same time.

Once the reconstruction has been completed by the stage II GPU, thealpha-correction process is initiated with the intention of filteringthe reconstructed image. This step performs a 2D separable box filterconvolution and a 2D separable comb filter convolution on thereconstructed data. The result is an image which has these periodicartifacts removed, which greatly improves the image quality. This setupmakes use of the RDMA1 capabilities of a professional grade GPU such asthe Nvidia Tesla K20c, and the faster processing capabilities of ahigh-end consumer grade GPU such as the Nvidia Titan Black.

In one embodiment of the present invention, a 160×80 configuration isused for multi-detector array 110 and 100×100 array of apertures 140used for collimation grid 90. An imaging frame can comprise 100×100detector images of 160×80 detector pixels. Each aperture 140 can act asa separate illumination point or source. Alternatively, the illuminationpoints or sources can be infrared, ultrasound, microwave, visible light,electromagnetic, gamma ray, proton or other form of radiation. Eachdetector image of 160×80 detector pixels can be a separate image of theobject. Since each aperture 140 is in a different position, the detectorimages of the object can be different for different apertures 140. The100×100 array of apertures 140 can be formed as 100 columns of 100 rowsor 100 rows of 100 columns. The imaging frame comprises of detectorimages from multiple illumination focal spot positions (in this example,100×100) with 160×80 detector pixels. In this example, the illuminationfocal spot positions are in a single plane. In an alternativeembodiment, the multiple illumination focal spot positions can belocated around the object.

Each imaging frame can be formed after a single scan of the electronbeam 40 over target 60 or 1307, dwelling on each targeted focal spot fora predetermined amount of time. Alternatively, the exposure time foreach focal spot can be broken into multiple illuminations (rescans)separated in time. Use of multiple illuminations can keep thetemperature of x-ray target 60 or 1307 below desired maximums (maximumsbeing set by the potential of the target to “burn out” when overheatedby the incident beam in a localized region). Thus, every focal-spotposition can be rescanned 2 or more times and the imaging frame formedfrom 2 or more scans of the electron beam over each focal spot in target60 or 1307.

A GPU with at least 3 Teraflops peak single-precision floating pointperformance can be used for GPU 1104. GPU 1104 can also act as apre-processing, buffer constitution, buffer construction or datarestructuring GPU operating as described herein. Memory bandwidth forboard with ECC off can be at least 200 Gigabytes per second. The RDMA1capability of GPU 1104 increases the data transfer rate from SDM 1102and SDM 1103. A GPU with at least 4 Teraflops peak single-precisionfloating point performance can be used for processing or reconstructionGPU 1105. Memory bandwidth can be at least 336 Gigabytes per second.

As electron beam 40 scans each targeted focal spot for a predeterminedamount of time, the imaging system stores a count of the number ofincoming photons striking the detector surface of each detector ofmulti-detector array 110 in the data buffer. In the exemplary embodimentwhere the exposure time for each focal spot is broken into multipleilluminations (rescans) separated in time, GPU 1104 can sum or aggregatethe photon counts for the illuminations or rescans over each targetedfocal spot for each detector of multi-detector array 110. In thismanner, an imaging frame can be developed with 100×100 detector imagesof 160×80 detector pixels. In an alternative embodiment, SDM 1102 andSDM 1103 or GPU 1104 can aggregate data from adjoining detector pixelelements and combine them into a single element with a lower bit valuethan the sum of 1-bit values of the adjoining detector pixels. In oneexample, four adjacent detector pixel elements with 1-bit values can becombined into a single element with a 3-bit value. With four adjacentdetector pixel elements, the maximum value (when all four detector pixelelements are 1 or ON) is 4, a 3-bit value. With a 160×80 configurationfor multi-detector array 110, a 160×80 element array with 1-bit valuescan be aggregate to an 80×40 element array with 3-bit values.

GPU 1104 can format and reshape the imaging frame and individualdetector images of the 100×100 detector images. GPU 1104 can store datafor each detector image of the 100×100 illumination focal spot positionscomprising 160×80 detector pixels into the proper address or location ofthe data buffer. GPU 1104 can also transpose the individual detectorimages. GPU 1104 can transform an imaging frame that is initially storedin row-by-row format (one row followed by another row, i.e. row major)into a column-by-column format (one column followed by another column,i.e. column major). GPU 1104 can also transform an imaging frame that isinitially stored in column-by-column format (one column followed byanother column, i.e. column major) into a row-by-row format (one rowfollowed by another row, i.e. row major). GPU 1104 can also verify thatthere are no dropped holes representing dropped detector images for theset of 100×100 detector images. Dropped holes or dropped detector imagescan occur if GPU 1104 cannot process data quickly enough to keep up withSDM 1102 and SDM 1103. GPU 1104 checks for dropped holes or droppeddetector images and reports it to the CPU. After completion ofpre-processing of the imaging frame, the aggregated and pre-processedimaging frame data can be distributed from GPU 1104 as a completepackage to processing or reconstruction GPU 1105 using a single burstmemory operation.

Under an alternative embodiment, GPU 1104 is coupled to and receivesdata from two additional SDMs in addition to SDM 1102 and SDM 1103. Eachpair of SDMs is coupled to and receives data from one scan datacontroller (SDC) 1101.

Under an alternative embodiment with 2 to N processing or reconstructionGPUs, GPU 1104 can split up the processing or reconstruction of animaging frame of 100×100 detector images of 160×80 detector pixels intosmaller imaging frames e.g. 10×10 detector images of 160×80 detectorpixels or 100×2 detector images of 160×80 detector pixels. The number ofdetector images in the smaller imaging frame and the number ofassociated apertures 140 can be 2.5%, 5%, 10%, 20%, 25%, 33%, 40%, 50%,60%, 70%, 75% or any percentage or range of percentages between 1% and99% of the total number of detector images or associated apertures 140.Smaller imaging frames can allow more processing or reconstruction GPUsto operate concurrently or can allow the task to be split up moreefficiently within a single GPU. Smaller imaging frames can also allowan individual reconstruction with involving that imaging frame to beaccomplished faster. However, management of reconstruction process isless complex with larger imaging frames and associated smaller quantityof imaging frames. A range between 5% and 10%, 10% and 20% or 25% and33% can be optimal.

The smaller imaging frames can comprise detector images for completerows, e.g. 100×1, 100×2, 100×3, 100×4 or 100×5, or complete columns,e.g. 1×100, 2×100, 3×100, 4×100 or 5×100, of apertures 140. Imagingframes with complete rows or complete columns can be beneficial becauseit can parallel the scan motion of the scanning beam through array ofapertures 140 reducing latency and possible conflict between detectordata flow and processing or reconstruction by GPU 1105. The smallerimaging frame can comprise detector images for 1%, 2%, 3%, 4%, 5%, 7%,10%, 15%, 20% or 25% of the total rows or columns or any number of rowsor columns or any range between 2% and 25%. Smaller imaging frames canallow processing or reconstruction to commence with less waiting timefor the scanning beam since the associated apertures 140 can be scannedfaster with a smaller number. Smaller imaging frames can also allow anindividual reconstruction with involving that imaging frame to beaccomplished faster. However, management of reconstruction process isless complex with larger imaging frames and associated smaller quantityof imaging frames. A range between 2% and 5% of the total rows orcolumns, 5% and 10% of the total rows or columns or 10% and 20% of thetotal rows or columns can be optimal.

Each of the smaller imaging frames comprising a different subset ornumber of the total number of illumination focal spot positions can betransmitted to different processing or reconstruction GPUs forprocessing or reconstruction. The processing or reconstruction GPUs canperform reconstruction with the smaller imaging frame using a priorlarger imaging frame e.g. 100×100 detector images of 160×80 detectorpixels, in effect using the smaller imaging frame to update the priorlarger imaging frame.

The foregoing descriptions of specific embodiments of the presentinvention have been presented for purposes of illustration anddescription. They are not intended to be exhaustive or to limit theinvention to the precise forms disclosed, and many modifications andvariations are possible in light of the above teaching. The embodimentswere chosen and described in order to best explain the principles of theinvention and its practical application, to thereby enable othersskilled in the art to best utilize the invention and various embodimentswith various modifications as are suited to the particular usecontemplated. It is intended that the scope of the invention be definedby the claims appended hereto and their equivalents.

What is claimed is:
 1. An x-ray imaging system for imaging an objectcomprising: a plurality of x-ray illumination source positions forproducing x-ray radiation at each of said illumination source positionsand projecting said x-ray radiation towards said object; a detector fordetecting said x-ray radiation from said object and transmittingdetector images for each of said illumination source positions; a memorybuffer for storing said detector images from said detector; a firstgraphics processing unit coupled to said memory buffer for formattingsaid detector images and for constructing a complete frame data set withsaid detector images for each of said illumination source positionsstored in said memory buffer; a second graphics processing unit coupledto said first graphics processing unit for receiving, from said firstgraphics processing unit, said complete frame data set with saiddetector images for each of said illumination source positions and forperforming image reconstruction on said complete frame data set; and abus coupled to said first graphics processing unit and said secondgraphics processing unit, wherein said complete frame data set with saiddetector images for each of said illumination source positions istransferred from said first graphics processing unit to said secondgraphics processing unit in a single burst operation.
 2. The imagingsystem of claim 1 further comprising: a scanning beam x-ray source. 3.The imaging system of claim 1 further comprising: a direct memory accessunit coupled to said first graphics processing unit for direct memoryaccess of memory by said first graphics processing unit.
 4. The imagingsystem of claim 1 further comprising: a direct memory access unitcoupled to said first graphics processing unit for direct memory accessof memory by said first graphics processing unit.
 5. The imaging systemof claim 1 further comprising: a direct memory access unit coupled tosaid second graphics processing unit for direct memory access of memoryby said second graphics processing unit.
 6. The imaging system of claim1 further comprising: a PCI-E bus coupled to said first graphicsprocessing unit and said second graphics processing unit.
 7. The imagingsystem of claim 1 further comprising: a third graphics processing unitcoupled to said first graphics processing unit and said memory bufferfor receiving said complete frame data set with said detector images foreach of said illumination source positions and performing imagereconstruction on said complete frame data set.
 8. The imaging system ofclaim 1 wherein said second graphics processing unit further comprises aplurality of blocks of worker threads capable of performing independentcalculations and memory operations.
 9. The imaging system of claim 1wherein said detector comprises a multi-detector array comprising aplurality of detector elements, wherein said memory buffer stores acount of incoming photons striking a surface of each of said detectorelements, wherein exposures of each focal spot in said multi-detectorarray are separated in time, and wherein said first graphics processingunit sums said count for said exposures over said each focal spot foreach of said detector elements.
 10. The imaging system of claim 1wherein said first graphics processing unit aggregates data fromadjoining pixel elements of said detector and combines them as a singleelement.
 11. The imaging system of claim 1 wherein said formatting saiddetector images comprises an operation selected from the group of:reshaping said detector images; transposing said detector images;transforming an image frame stored in a row-by-row format into acolumn-by-column format; and transforming an image frame stored in acolumn-by-column format into a row-by-row format.
 12. A method for x-rayimaging of an object comprising: producing x-ray radiation from aplurality of x-ray illumination source positions; directing said x-rayradiation towards said object; measuring an amount of said x-rayradiation striking a detector; producing detector images of said objectfor each of said plurality of x-ray illumination source positions basedon said amount of said x-ray radiation striking said detector; using afirst graphics processing unit to construct a complete frame data setwith said detector images for each of said illumination sourcepositions; transmitting said complete frame data set with said detectorimages for each of said illumination source positions from said firstgraphics processing unit to a second graphics processing unit in asingle burst operation; and using said second graphics processing unitto perform image reconstruction on said complete frame data set receivedfrom said first graphics processing unit.
 13. The method of claim 12further comprising: producing said complete frame data set at least onceevery 130 milliseconds.
 14. The method of claim 12 further comprising:identifying any of said detector images that are dropped from saidcomplete frame data set using said first graphics processing unit. 15.The method of claim 12 further comprising: aggregating image data foreach of said illumination source positions to form said detector images.16. The method of claim 12 further comprising: transferring saidcomplete frame data set with said detector images for each of saidillumination source positions from said first graphics processing unitto said second graphics processing unit at a rate of 30 Hz.
 17. Themethod of claim 12 further comprising: organizing computation for saidimage reconstruction in blocks of worker threads wherein each of saidworker threads performs independent calculations and memory operations.18. The method of claim 12 further comprising: aggregating image datafor each of said illumination source positions to form said detectorimages; and organizing computation for said image reconstruction inblocks of worker threads wherein each of said worker threads performsindependent calculations and memory operations.
 19. The method of claim12 further comprising: transmitting detector images from said detectorat a rate of at least 1 MHz.
 20. The method of claim 12 furthercomprising: transposing said complete frame data set using said firstgraphics processing unit.